1. Field of the Invention
The present invention concerns a method for the generation of an MR image of a vascular structure of an examination region, and an MR apparatus for this purpose. In the following, the term “angiography image” shall be used in the generic sense; it does not necessarily encompass only a single two-dimensional image, but depending on the context, may also encompass an angiographic image data set of a desired volume.
2. Description of the Prior Art
In the field of MR angiography, methods that function without the need for contrast agents are gaining significance. One class of such methods exploits the pulsatile nature of the arterial blood flow. With signal recording (data acquisition) of a flow-sensitive sequence at the point in time of the quicker flow rate (systole), one obtains, in the ideal case, a deletion (absence) of the arterial signal. With a signal recording at the point in time in the cardiac cycle at which point the rate is at a minimum, or zero (diastole), one obtains the arteries ideally with a bright signal. Through the subtraction of the two recorded data sets, the signal of the stationary tissue can be removed, so an arterial angiography is obtained. While very good results have been obtained with this method in healthy subjects, the results for other patients leaves something to be desired. Particularly with the presence of pathologies, the pulsatile dynamics may be only very weakly pronounced, or even at a complete standstill, such that only a limited difference, or none at all, occurs between the blood flow rate during systole and the blood flow rate in diastole.
Another class of such methods uses a marking of the examination region in question for suppressing the statistic tissue signal (so-called spin labeling). During a waiting period tin, fresh unmarked blood can flow in from externally, and this blood is brightly depicted by means of the following signal acquisition. Generally, an additional measure for suppression is provided, for example by means of the upstream activation of a fat saturation module. Usually it is also advantageous to configure the measurement to the heartbeat and thereby to the flow characteristic, i.e. to place the marking temporally prior to the systolic pulse wave, and to locate the waiting period during the pulse wave, and the signal recording directly after the pulse wave. The marking, or the so-called labeling, can be implemented, for example, as layer-selective saturation or inversion, and conventionally, the imaging sequence is implemented as a three-dimensional procedure. A method of this type is described, for example, by Wyttenbach et al. (Renal artery assessment with non-enhanced steady-state free precession versus contrast-enhanced MR angiography. Radiology 2007; 245(1): 186-195). A fundamental problem with this method is that the vessels of the volume in question must be filled during tin with fresh blood, flowing in from outside of the marked volume. Therefore, 3D methods with larger marked volumes are more suited for regions with very fast or wider-reaching blood flows, e.g. in the head, the aorta or the kidney arteries. In the peripheral regions of the body the blood remains still during the major portion of the cardiac cycle, and is then quickly pushed forward during the pulse wave. Even with healthy persons, the distance traveled amounts to only a few centimeters, such that in the legs, hardly any long visual field can be used extending in the direction of the legs, designated herein as the z-axis. This limitation is more striking with patients whose pulsatile dynamics are less pronounced, wherein the distance traveled by the blood during the remaining weak pulse wave may be significantly shorter.
In Edelman et al., in Magn Reson Med 63:951-958 (2010), with the title, “Quiescent-Interval Single-Shot Unenhanced Magnetic Resonance Angiography of Peripheral Vascular Disease: Technical Considerations and Clinical Feasibility,” the use of a 2D method is described, in which the blood must only be transported by the pulse wave over the thickness of a two-dimensional layer, which is reliably fulfilled even with unhealthy patients. A single axial layer is thereby recorded with each heartbeat. Typically, layer thicknesses of 3 mm and 1.4 mm are used. To create thinner thicknesses, an extension of the RF pulse, and thus the repetition periods TR, is required, which is undesired. Furthermore, with even thinner layers, the obtainable signal-to-noise ratio is marginally low, in particular with the clinically, currently most commonly used, field strengths of 1.5 tesla. This spatial resolution, limited in the z-axis, represents a major limitation of the method, in particular when finer vessels do not run strictly along the z-axis, as is the case, for example, in regions of the trifurcation in the lower leg, or with pathologies.
Another (as yet not used) alternative could be a three-dimensional imaging with very thin partitions. A problem with this alternative would be that efficiency would not be increased because the time for encoding a 3D partition is the same as that for the acquisition of a 2D layer. To improve the resolution on the z-axis from, for example, 3 mm to 0.75 mm, the measuring time must be increased by a factor of 4 accordingly. This is a problem, because, for example, with an angiography of the pelvis-leg region, normally a coverage from the feet up to the abdominal region is desired.